Magnetic resonance system and method employing a digital squid

ABSTRACT

A magnetic resonance system, comprising at least one SQUID, configured to receive a radio frequency electromagnetic signal, in a circuit configured to produce a pulsatile output having a minimum pulse frequency of at least 1 GHz which is analyzed in a processor with respect to a timebase, to generate a digital signal representing magnetic resonance information. The processor may comprise at least one rapid single flux quantum circuit. The magnetic resonance information may be image information. A plurality of SQUIDs may be provided, fed by a plurality of antennas in a spatial array, to provide parallel data acquisition. A broadband excitation may be provided to address a range of voxels per excitation cycle. The processor may digitally compensate for magnetic field inhomogeneities.

CROSS REFERENCE TO RELATED APPLICATIONS

The present application claims benefit of priority from U.S. ProvisionalPatent Application Ser. No. 61/264,032, filed Nov. 24, 2009, theentirety of which is expressly incorporated herein by reference.

BACKGROUND OF THE INVENTION

Magnetic resonance imaging (MRI) is an established technology for highresolution three-dimensional imaging of the living human body, withoutthe need for potentially hazardous x-rays. See, Joseph P. Hornak, Ph.D.,The Basics of MRI, [[http://]]www.cis.rit.edu/htbooks/mri/ (copyright1996-2010), expressly incorporated herein by reference. This is based onmagnetic resonance of the magnetic moments of hydrogen nuclei (protons)in the presence of a magnetic field. The process involves resonantabsorption and emission of a radio-frequency (RF) signal in the tissue,at a frequency given by f=γB, where B is the field in teslas (T) andγ=43 MHz/T for the proton. Since the signal strength and the spatialresolution generally increase with larger fields, conventional MRItechnology typically uses a field of at least B=1.5 T, corresponding tof=65 MHz. A uniform, low-noise magnetic field of this magnitude requiresa superconducting magnet, which yields a large and bulky system.

MRI is sensitive to the magnetic field gradient, temperature, andmicroenvironment. By using a strong, uniform field, many relatively weakcontributing factors are generally ignored in the analysis. However,using powerful computing devices, it is possible to analyze each voxel(minimum unit of volume in the image space) dependent on its particularmagnetic field strength, gradient, and therefore draw conclusions aboutits microenvironment and structure, which is the typical goal of an MRIprocedure.

One disadvantage of this conventional method is that it can be quiteslow to achieve high resolution imaging over a substantial volume oftissue. An imaging time of many minutes can be inconvenient for thepatient, as well as being inconsistent with moving tissues such as theheart and lungs. Functional MRI (fMRI) acquires a series of MRI imageswhich are time-synchronized with a physiological status, i.e.,heartbeat, to permit imaging of moving tissues, and perhaps moreimportantly, an analysis of their dynamic attributes.

The slow imaging time is limited by scanning magnetic field gradients inthree dimensions to select a small region corresponding to the resonancecondition. A single receiver coil typically measures this RF signal. Oneapproach to accelerating the image acquisition is to use a plurality ofreceiver coils, with different locations, in order to obtain somespatial information and reduce the need for the full gradient scanning.A fully parallel approach which can eliminate entirely at least onedimension of gradient scanning is sometimes called “single echoacquisition”, and comprises using an array of N closed spaced (butdecoupled) receiver coils, where N=64 coils has been demonstrated. (See,for example, U.S. Pat. No. 6,771,071, “Magnetic resonance imaging usinga reduced number of echo acquisitions”, expressly incorporated herein byreference.) While this does increase the image acquisition speed, thisapproach has not yet been adopted for commercial MRI systems, due inpart to the complex network of RF receivers required.

A recently developed alternative technology for MRI uses ultra-lowmagnetic fields, with measurement fields as low as the microtesla range,and frequencies as low as 1 kHz instead of 65 MHz. See, for example,U.S. Pat. Nos. 6,159,444; 6,885,192; 7,053,410; 7,117,102; 7,218,104;7,187,169; and published US patent application 2009/0072828, expresslyincorporated herein by reference.

This runs counter to the conventional trend of using ever highermagnetic fields to improve signal integrity and resolution. However,this approach can make use of ultra-sensitive magnetic field detectorscalled superconducting quantum interference devices, or SQUIDs. Thesensitivity of the SQUID detector can partially compensate for the weaksignal at these very low frequencies. A system of this type does notaddress the imaging speed issue (which may even be slower due toadditional required signal averaging), but such a system would besubstantially more compact and portable due to the small magnetic field,and therefore likely small magnet. Still, work to date has demonstratedonly a fairly low-resolution system, which would limit itsapplicability.

One suggested improvement on the ultra-low field MRI system, to increasethe signal-to-noise ratio and the resolution, is to use a plurality ofSQUID receivers. This would also enable the same system to be used formagnetoencephalography (MEG) or magnetocardiography (MCG), where thearray of SQUIDs could also localize sources of electrical currents innerves or heart muscle. (See, for example, U.S. Pat. No. 7,573,268,“Direct imaging of neural currents using ultralow field magneticresonance techniques”, expressly incorporated herein by reference, andV. Zotev et al, “Parallel MRI at Microtesla Fields”, J. MagneticResonance, vol. 192, p. 197, 2008.) However, this partially parallelapproach did not address how this small array (with 7 elements) could bescaled to much larger numbers to achieve large scan-time accelerationfactors.

It is well known in the art that SQUIDs are capable of ultra-sensitivemeasurements of magnetic fields at low frequencies of order a kilohertzor less. However, it is less well known that the SQUID itself isactually a high-frequency device, capable of field measurement up to GHzfrequencies. The frequency limitation in conventional SQUID systems isactually in the external control loop that extends the dynamic range andlinearity of the device as an analog sensor of magnetic field strengthand/or gradient. Two complementary approaches have been developed toadapt SQUIDs to practical RF systems. In one approach, arrays of SQUIDs(sometimes called superconducting quantum interference filters or SQIFs)are coupled together to increase linearity and dynamic range, withoutrequiring an external control loop. (See, for example, U.S. Pat. Nos.6,690,162 and 7,369,093, expressly incorporated herein by reference.) Inanother approach, sometimes called a digital SQUID or a SQUID-baseddigitizer, the SQUID is used to generate fast single-flux-quantum (SFQ)voltage pulses, which are processed by rapid-single-flux-quantum (RSFQ)superconducting digital logic circuits to achieve a combination oflinearity and dynamic range, as well as the flexibility of digitalprocessing. See, for example, U.S. Pat. Nos. 5,420,586; 7,365,663;7,598,897, expressly incorporated herein by reference.

The present technology provides improvements for an MRI system that mayhelp achieve the dual goals of fine imaging, while also obtainingsubstantial acceleration of image acquisition. Likewise, the presenttechnology may also assist in reducing the system bulk, and enhancingflexibility.

A useful background of MRI is found in[http://]en.wikipedia.org/wiki/Magnetic_resonance_imaging (Nov. 22,2009), expressly incorporated herein by reference. In order to acquire aMagnetic Resonance Image, Radio frequency (RF) fields are used toperiodically align a magnetic moment of a portion of hydrogen (orcertain other atomic isotopes), which then relax to their unalignedstate over time. Certain nuclei such as ¹H (protons), ²H, ³He, ¹³C, ²³Naor ³¹P, have a non-zero spin and therefore a magnetic moment. In thecase of the so-called spin-½ nuclei, such as ¹H, there are two spinstates, sometimes referred to as “up” and “down”. When these spins areplaced in a strong external magnetic field they precess around an axisalong the direction of the field. Protons align in two energy“eigenstates” (the “Zeeman effect”): one low-energy and one high-energy,which are separated by a very small splitting energy.

An image may be made, on a per-voxel basis, of the magnitude andrelaxation time of the magnetic alignment. The frequency at which theprotons resonate depends on the strength of the magnetic field. Thisfield-strength dependence therefore allows a frequency encoding ofposition. By superposing fields which predictably alter the magneticfield in known manner, different coordinates may be “addressed”,allowing full image formation, for example as a set of slices, or as athree dimensional matrix.

Different tissues can be distinguished because different chemicals,prototypically water and lipid, can be detected because the protons indifferent chemical compositions return to their equilibrium state atdifferent rates. There are other aspects of magnetic resonance that canalso be exploited to extract information. For example, in addition torelaxation time, local environments can also create perturbation ofmagnetic fields, and the presence of characteristic perturbations can beused to infer the local environment. For example, a single proton canpredictably perturb the magnetic field of another proton locatedproximately on the same molecule.

In the static magnetic fields commonly used in MRI, the energydifference between the nuclear spin states corresponds to a radiofrequency photon. Resonant absorption of energy by the protons due to anexternal oscillating magnetic field will occur at the Larmor frequencyfor the particular nucleus.

The net magnetization vector has two states. The longitudinalmagnetization is due to a tiny excess of protons in the lower energystate. This gives a net polarization parallel to the external field.Application of an RF pulse can tip sideways (with i.e. a so-called 90°pulse) or even reverse (with a so-called 180° pulse) this netpolarization vector.

The recovery of longitudinal magnetization is called longitudinal or T₁relaxation and occurs exponentially with a time constant T₁. The loss ofphase coherence in the transverse plane is called transverse or T₂relaxation. T₁ is thus associated with the enthalpy of the spin system(the number of nuclei with parallel versus anti-parallel spin) while T₂is associated with its entropy (the number of nuclei in phase).

When the radio frequency pulse is turned off, the transverse vectorcomponent produces an oscillating magnetic field which induces a smallcurrent in the receiver coil. This signal is called the free inductiondecay (FID). In an idealized nuclear magnetic resonance experiment, theFID decays approximately exponentially with a time constant T₂, but inpractical MRI small differences in the static magnetic field atdifferent spatial locations (“inhomogeneities”) cause the Larmorfrequency to vary across the body creating destructive interferencewhich shortens the FID. The time constant for the observed decay of theFID is called the T₂* relaxation time, and is always shorter than T₂.Also, when the radio frequency pulse is turned off, the longitudinalmagnetization starts to recover exponentially with a time constant T₁.

In MRI, the static magnetic field is caused to vary across the body (afield gradient), so that different spatial locations become associatedwith different precession frequencies. Usually these field gradients arepulsed, and a variety of RF and gradient pulse sequences may be used.Application of field gradient destroys the FID signal, but this can berecovered and measured by a refocusing gradient (to create a so-called“gradient echo”), or by a radio frequency pulse (to create a so-called“spin-echo”). The whole process can be repeated when some T₁-relaxationhas occurred and the thermal equilibrium of the spins has been more orless restored.

Typically in soft tissues T₁ is around one second while T₂ and T₂* are afew tens of milliseconds, but these values vary widely between differenttissues (and different external magnetic fields), permitting MRIdistinguish different types of soft tissues. Contrast agents work byaltering (shortening) the relaxation parameters, especially T₁.

A number of schemes have been devised for combining field gradients andradio frequency excitation to create an image: 2D or 3D reconstructionfrom projections, much as in Computed Tomography; Building the imagepoint-by-point or line-by-line; Gradients in the RF field rather thanthe static field. Although each of these schemes is occasionally used inspecialist applications, the majority of MR Images today are createdeither by the Two-Dimensional Fourier Transform (2DFT) technique withslice selection, or by the Three-Dimensional Fourier Transform (3DFT)technique. Another name for 2DFT is spin-warp. The 3DFT technique israther similar except that there is no slice selection andphase-encoding is performed in two separate directions. Another schemewhich is sometimes used, especially in brain scanning or where imagesare needed very rapidly (such as in functional MRI or fMRI), is calledecho-planar imaging (EPI): In this case, each RF excitation is followedby a train of gradient echoes with different spatial encoding.

Image contrast is created by differences in the strength of the NMRsignal recovered from different locations within the sample. Thisdepends upon the relative density of excited nuclei (usually waterprotons), on differences in relaxation times (T₁, T₂, and T₂*) of thosenuclei after the pulse sequence, and often on other parameters. Contrastin most MR images is actually a mixture of all these effects, butcareful design of the imaging pulse sequence allows one contrastmechanism to be emphasized while the others are minimized. In the brain,T₁-weighting causes the nerve connections of white matter to appearwhite, and the congregations of neurons of gray matter to appear gray,while cerebrospinal fluid (CSF) appears dark. The contrast of whitematter, gray matter and cerebrospinal fluid is reversed using T₂ or T₂*imaging, whereas proton-density-weighted imaging provides littlecontrast in healthy subjects. Additionally, functional parameters suchas cerebral blood flow (CBF), cerebral blood volume (CBV) or bloodoxygenation can affect T₁, T₂ and T₂* and so can be encoded withsuitable pulse sequences.

In some situations it is not possible to generate enough image contrastto adequately show the anatomy or pathology of interest by adjusting theimaging parameters alone, in which case a contrast agent may beadministered. This can be as simple as water, taken orally, for imagingthe stomach and small bowel. However, most contrast agents used in MRIare selected for their specific magnetic properties. Most commonly, aparamagnetic contrast agent (usually a gadolinium compound) is given.Gadolinium-enhanced tissues and fluids appear extremely bright onT₁-weighted images. This provides high sensitivity for detection ofvascular tissues (e.g., tumors) and permits assessment of brainperfusion (e.g., in stroke).

More recently, superparamagnetic contrast agents, e.g., iron oxidenanoparticles, have become available. These agents appear very dark onT₂-weighted images and may be used for liver imaging, as normal livertissue retains the agent, but abnormal areas (e.g., scars, tumors) donot. They can also be taken orally, to improve visualization of thegastrointestinal tract, and to prevent water in the gastrointestinaltract from obscuring other organs (e.g., the pancreas). Diamagneticagents such as barium sulfate have also been studied for potential usein the gastrointestinal tract, but are less frequently used.

In 1983 Ljunggren and Tweig independently introduced the k-spaceformalism, a technique that proved invaluable in unifying different MRimaging techniques. They showed that the demodulated MR signal S(t)generated by freely precessing nuclear spins in the presence of a linearmagnetic field gradient G equals the Fourier transform of the effectivespin density. Mathematically:S(t)={tilde over (ρ)}_(eff)({right arrow over (k)}(t))≡∫d{right arrowover (x)}ρ({right arrow over (x)}x)·e^(2πi{right arrow over (k)}(t)·{right arrow over (x)})

where:

{right arrow over (k)}(t)≡∫₀ ^(t){right arrow over (G)}(τ)dτ

In other words, as time progresses the signal traces out a trajectory ink-space with the velocity vector of the trajectory proportional to thevector of the applied magnetic field gradient. By the term effectivespin density we mean the true spin density ρ({right arrow over (x)})corrected for the effects of T₁ preparation, T₂ decay, dephasing due tofield inhomogeneity, flow, diffusion, etc., and any other phenomena thataffect that amount of transverse magnetization available to inducesignal in the RF probe. From the basic k-space formula, it followsimmediately that an image I({right arrow over (x)}) may be constructedby taking the inverse Fourier transform of the sampled data:I({right arrow over (x)})=∫d{right arrow over (x)}S({right arrow over(k)}(t))·e ^(−2πi{right arrow over (k)}(t)·{right arrow over (x)})

In a standard spin echo or gradient echo scan, where the readout (orview) gradient is constant (e.g. G), a single line of k-space is scannedper RF excitation. When the phase encoding gradient is zero, the linescanned is the k_(x) axis. When a non-zero phase-encoding pulse is addedin between the RF excitation and the commencement of the readoutgradient, this line moves up or down in k-space, i.e., we scan the linek_(y)=constant. In single-shot EPI, all of k-space is scanned in asingle shot, following either a sinusoidal or zig-zag trajectory. Sincealternating lines of k-space are scanned in opposite directions, thismust be taken into account in the reconstruction. Multi-shot EPI andfast spin echo techniques acquire only part of k-space per excitation.In each shot, a different interleaved segment is acquired, and the shotsare repeated until k-space is sufficiently well-covered. Since the dataat the center of k-space represent lower spatial frequencies than thedata at the edges of k-space, the T_(E) value for the center of k-spacedetermines the image's T₂ contrast.

The importance of the center of k-space in determining image contrastcan be exploited in more advanced imaging techniques. One such techniqueis spiral acquisition—a rotating magnetic field gradient is applied,causing the trajectory in k-space to spiral out from the center to theedge. Due to T₂ and T₂* decay the signal is greatest at the start of theacquisition, hence acquiring the center of k-space first improvescontrast to noise ratio (CNR) when compared to conventional zig-zagacquisitions, especially in the presence of rapid movement.

Since {right arrow over (x)} and {right arrow over (k)} are conjugatevariables (with respect to the Fourier transform) we can use the Nyquisttheorem to show that the step in k-space determines the field of view ofthe image (maximum frequency that is correctly sampled) and the maximumvalue of k sampled determines the resolution for each axis, i.e.,

${FOV} \propto {\frac{1}{\Delta\; k}\mspace{14mu}{Resolution}} \propto {{k_{\max}}.}$

In acquiring a typical MRI image, Radio frequencies are transmitted atthe Larmor frequency of the nuclide to be imaged. For example, for ¹H ina magnetic field of 1 T, a frequency of 42.5781 MHz would be employed.During the first part of the pulse sequence, a shaped pulse, e.g., usingsinc modulation, causes a 90° nutation of longitudinal nuclearmagnetization within a slab, or slice, creating transversemagnetization. During the second part of the pulse sequence, a phaseshift is imparted upon the slice-selected nuclear magnetization, varyingwith its location in the Y direction. During the third part of the pulsesequence, another slice selection (of the same slice) uses anothershaped pulse to cause a 180° rotation of transverse nuclearmagnetization within the slice. This transverse magnetization refocusesto form a spin echo at a time TE. During the spin echo, afrequency-encoding (FE) or readout gradient is applied, making theresonant frequency of the nuclear magnetization vary with its locationin the X direction. The signal is sampled n_(FE) times by the ADC duringthis period. Typically n_(FE) of between 128 and 512 samples are taken.The longitudinal magnetization is then allowed to recover somewhat andafter a time TR the whole sequence is repeated n_(PE) times, but withthe phase-encoding gradient incremented. Typically n_(PE) of between 128and 512 repetitions are made. Negative-going lobes in G_(X) and G_(Z)are imposed to ensure that, at time TE (the spin echo maximum), phaseonly encodes spatial location in the Y direction. Typically TE isbetween 5 ms and 100 ms, while TR is between 100 ms and 2000 ms.

After the two-dimensional matrix (typical dimension between 128×128 and512×512) has been acquired, producing the so-called K-space data, atwo-dimensional Fourier transform is performed to provide the familiarMR image. Either the magnitude or phase of the Fourier transform can betaken, the former being far more common.

The magnet is a large and expensive component of an MRI scanner. Thestrength of the magnet is measured in tesla (T). Clinical magnetsgenerally have a field strength in the range 0.1-3.0 T, with researchsystems available up to 9.4 T for human use and 21 T for animal systems.Just as important as the strength of the main magnet is its precision.The straightness of the magnetic lines within the center (or, as it istechnically known, the iso-center) of the magnet needs to benear-perfect. This is known as homogeneity. Fluctuations(inhomogeneities in the field strength) within the scan region should beless than three parts per million (3 ppm). Three types of magnets havebeen used:

Permanent magnet: Conventional magnets made from ferromagnetic materials(e.g., iron alloys or compounds containing rare earth elements such asneodymium) can be used to provide the static magnetic field. A permanentmagnet that is powerful enough to be used in a traditional type of MRIwill be extremely large and bulky; they can weigh over 100 tons.Permanent magnet MRIs are very inexpensive to maintain; this cannot besaid of the other types of MRI magnets, but there are significantdrawbacks to using permanent magnets. They are only capable of achievingweak field strengths compared to other MRI magnets (usually less than0.4 T) and they are of limited precision and stability. Permanentmagnets also present special safety issues; since their magnetic fieldscannot be “turned off,” ferromagnetic objects are virtually impossibleto remove from them once they come into direct contact. Permanentmagnets also require special care when they are being brought to theirsite of installation.

Resistive electromagnet: A solenoid wound from copper wire is analternative to a permanent magnet. An advantage is low initial cost, butfield strength and stability are limited. The electromagnet requiresconsiderable electrical energy during operation which can make itexpensive to operate. This design is considered obsolete for typicalapplication.

Superconducting electromagnet: When a niobium-titanium or niobium-tinalloy is cooled by liquid helium to 4 K (−269° C., −452° F.) it becomesa superconductor, losing resistance to flow of electrical current. Anelectromagnet constructed with superconductors can have extremely highfield strengths, with very high stability. The construction of suchmagnets is costly, and the cryogenic helium imposes operating costs andrequires careful handling. However, despite their cost, helium cooledsuperconducting magnets are the most common type found in MRI scannerstoday.

Magnetic field strength is an important factor in determining imagequality. Higher magnetic fields increase signal-to-noise ratio,permitting higher resolution or faster scanning. However, higher fieldstrengths require more costly magnets with higher maintenance costs, andhave increased safety concerns. A field strength of 1.0-1.5 T isconsidered a good compromise between cost and performance for generalmedical use. However, for certain specialist uses (e.g., brain imaging)higher field strengths are desirable, with some hospitals now using 3.0T scanners.

Gradient coils are used to spatially encode the positions of protons byvarying the magnetic field linearly across the imaging volume. TheLarmor frequency will then vary as a function of position in the x, yand z-axes. Gradient coils are usually resistive electromagnets poweredwhich permit rapid and precise adjustments to their field strength anddirection. Typical gradient systems are capable of producing gradientsfrom 20 mT/m to 100 mT/m (i.e., in a 1.5 T magnet, when a maximal Z-axisgradient is applied, the field strength may be 1.45 T at one end of a 1m long bore and 1.55 T at the other). It is the magnetic gradients thatdetermine the plane of imaging—because the orthogonal gradients can becombined freely, any plane can be selected for imaging. Scan speed istypically dependent on performance of the gradient system. Strongergradients allow for faster imaging, or for higher resolution; similarly,gradients systems capable of faster switching can also permit fasterscanning. However, gradient performance may also be limited by safetyconcerns over possible nerve stimulation.

The radio frequency (RF) transmission system consists of an RFsynthesizer, power amplifier and transmitting coil. The receiver in atraditional system consists of a receiving coil, pre-amplifier andsignal processing system. It is possible to scan using an integratedcoil for RF transmission and MR signal reception, but if a small regionis being imaged, then better image quality (i.e. higher signal-to-noiseratio) is obtained by using a close-fitting smaller coil.

A recent development in MRI technology has been the development ofsophisticated multi-element phased array coils which are capable ofacquiring multiple channels of data in parallel. See, Roemer P B,Edelstein W A, Hayes C E, Souza S P, Mueller O M (1990). “The NMR phasedarray”. Magnetic Resonance in Medicine 16 (2): 192-225. This ‘parallelimaging’ technique uses acquisition schemes that allow for acceleratedimaging, by replacing some of the spatial coding originating from themagnetic gradients with the spatial sensitivity of the different coilelements. However, the increased acceleration also reduces thesignal-to-noise ratio and can create residual artifacts in the imagereconstruction. Two frequently used parallel acquisition andreconstruction schemes are known as SENSE and GRAPPA. See, Pruessmann KP, Weiger M, Scheidegger M B, Boesiger P (1999). “SENSE: Sensitivityencoding for fast MRI”. Magnetic Resonance in Medicine 42 (5): 952-962.doi:10.1002/(SICI)1522-2594(199911)42:5<952::AID-MRM16>3.0.00; 2-S. PMID10542355, Griswold M A, Jakob P M, Heidemann R M, Nittka M, Jellus V,Wang J, Kiefer B, Haase A (2002). “Generalized autocalibrating partiallyparallel acquisitions (GRAPPA)”. Magnetic Resonance in Medicine 47 (6):1202-1210.

doi:10.1002/mrm.10171. PMID 12111967; Blaimer M, Breuer F, Mueller M,Heidemann R M, Griswold M A, Jakob P M (2004). “SMASH, SENSE, PILS,GRAPPA: How to Choose the Optimal Method”. Topics in Magnetic ResonanceImaging 15 (4): 223-236.[http://]cfmriweb.ucsd.edu/ttliu/be280a_(—)05/blaimer05.pdf.

Hydrogen is the most frequently imaged nucleus in MRI because it ispresent in biological tissues in great abundance. However, any nucleuswhich has a net nuclear spin could potentially be imaged with MRI. Suchnuclei include helium-3, carbon-13, fluorine-19, oxygen-17, sodium-23,phosphorus-31 and xenon-129. ²³Na and ³¹P are naturally abundant in thebody, so can be imaged directly. Gaseous isotopes such as ³He or ¹²⁹Xemust be hyperpolarized and then inhaled as their nuclear density is toolow to yield a useful signal under normal conditions. ¹⁷O, ¹³C and ¹⁹Fcan be administered in sufficient quantities in liquid form (e.g.¹⁷O-water, ¹³C-glucose solutions or perfluorocarbons) thathyperpolarization is not a necessity.

Multinuclear imaging is primarily a research technique at present.However, potential applications include functional imaging and imagingof organs poorly seen on ¹H MRI (e.g. lungs and bones) or as alternativecontrast agents. Inhaled hyperpolarized ³He can be used to image thedistribution of air spaces within the lungs. Injectable solutionscontaining ¹³C or stabilized bubbles of hyperpolarized ¹²⁹Xe have beenstudied as contrast agents for angiography and perfusion imaging. ³¹Pcan potentially provide information on bone density and structure, aswell as functional imaging of the brain.

Portable magnetic resonance instruments are available for use ineducation and field research. Using the principles of Earth's field NMR,they have no powerful polarizing magnet, so that such instruments can besmall and inexpensive. Some can be used for both EFNMR spectroscopy andMRI imaging, e.g., the Terranova-MRI Earth's Field MRI teaching system([http://]www.magritek.com/terranova.html). The low strength of theEarth's field results in poor signal to noise ratios, requiring longscan times to capture spectroscopic data or build up MRI images.

A magnetic resonance system that can flexibly combine a plurality ofsignals and antennas, permitting faster generation of high-resolutionmedical images in a safe and economical system, is desirable.

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SUMMARY OF THE INVENTION

A system and method for radio reception for magnetic resonance imagingis disclosed. According to one embodiment, direct digitization of theradio frequency signal, without frequency shifting, using at least oneJosephson junction to generate a series of single-flux-quantum voltagepulses at a rate which is much larger than the frequency of the radiosignal, is employed. These pulses are digitally processed to generate adigital representation of the radio frequency signal.

In one embodiment, this receiver can receive a wideband signal thatcomprises a plurality of narrowband signals such as are conventionallyreceived in known MRI systems. These narrowband signals may be receivedfrom a plurality of input antennas, wherein the signals may be subjectto analog combination before digitization. Alternatively, digitalsignals from a plurality of antennas may be digitally multiplexed afterdirect digitization. This system and method enables several techniquesthat lead to faster imaging with improved spatial resolution.

According to a preferred embodiment, the RF sensors are superconductingmagnetic field sensors, e.g., SQUIDs, and are provided in a spatialarray, to therefore provide spatial encoding of magnetic field data. TheSQUIDs preferably respond to magnetic fields up to high frequencies. Forexample, for hydrogen in a 10 T magnetic field, a sensor capable ofdetecting a 400 MHz signal is required. As a digital sensor, a degree ofoversampling is appropriate, e.g., at least 2 times, and preferably atleast 4-8 times. For lower magnetic field strength, the characteristicfrequency is lower.

One aspect of the system and method therefore provides a radio frequencyreceiver for Magnetic Resonance Imaging using Single-Flux-QuantumDigital Electronics. This system preferably comprises at least onedevice that converts a radio-frequency magnetic flux signal to anoversampled digital representation. The device preferably comprises atleast one Josephson junction. The oversampled digital representation maycomprise a series of single-flux-quantum voltage pulses. An oversampledrepresentation may be reduced in sampling rate to obtain additional bitsof precision.

In one embodiment of the invention, the RF magnetic flux digitizercomprises a superconducting analog-to-digital converter (ADC), asdescribed in “Superconducting analog-to-digital converters”, O. Mukhanovet al., Proceedings of the IEEE, vol. 92, p. 1564, 2004. For example,this ADC may comprise a phase-modulation-demodulation ADC which issampled at a clock rate in excess of 1 GHz, and which exhibits aflux-equivalent noise spectral density of less than about 10 μΦ₀/√Hz,where Φ₀=h/2e=2 fT-m² is the magnetic flux quantum. Such an ADCdemonstrates both the sensitivity and the linear dynamic range neededfor digitizing the weak flux signal present in MRI measurements. In thecontext of this invention, such a superconductor ADC with magnetic fluxinput provides an example of an integrated digital SQUID.

A traditional MRI system downconverts a narrowband signal from amagnetic coil sensor, producing a baseband signal which is thendigitized. On the other hand, one embodiment of the present systememploys a large bandwidth high frequency digitizer that operatesdirectly on the radio frequency signal. Further, the sensor is able todetect small perturbations when operating within relatively largemagnetic fields, i.e., it exhibits a large dynamic range.

While one embodiment of the system employs an array of sensors, anotherembodiment employs a more traditional architecture with a single pickupcoil, with a wideband detector.

The present technology provides, in accordance with one embodiment, anMRI system which comprises an array of receiver coils, each associatedwith a SQUID-based magnetic field sensor. In a preferred embodiment, asshown in FIG. 1, there are N receivers, where N may be a large number,for example, 100-1000. All receivers may be mounted together in the samecryostat, for convenience of refrigeration. Each receiver comprises acoupling coil and a SQUID-based transducer and digitizer.

The coupling coil may be a traditional conductive wire coil, or anintegrated inductive coil on a common substrate with the SQUID, forexample. The coupling coil may comprise a narrowband resonant antennawith capacitance as well as inductance, or alternatively may be abroadband antenna. In some embodiments, the coupling coil may comprise asuperconducting material; in others it may comprise a low-resistivitymetal such as copper. According to the present superconductortechnology, a superconducting coil and SQUID must be cooled to anappropriate cryogenic temperature for proper operation. However, even aresistive coil will generally reduce its resistance and exhibit improvedlow-noise performance at cryogenic temperatures. In one embodiment, theflux signal from the coupling coil is fed directly to the SQUID. In analternative embodiment, the flux signal may first be amplified using anappropriate low-noise amplifier (LNA), which may comprise asemiconductor device, and may be cooled to cryogenic temperatures toreduce noise. The output of this LNA would then be coupled to the SQUID.

The digitizer may generate an oversampled representation of the RFsignal, which is digitally downconverted and averaged to obtain thebaseband (envelope) signal that is used for imaging. The resultingmulti-bit digital signal may be serialized and multiplexed with similarsignals (e.g., to reduce the number of output lines from the cryostat),then sent to a digital processing module at room temperature.

The SQUIDs may be low temperature superconductors, i.e., operating below20 K, high temperature superconductors, operating below −70° C. (˜200K), and, as they emerge, room temperature superconductors. Preferably,according to currently available devices, low temperaturesuperconductors are employed. While such cryogenic cooling can beobtained using a liquid cryogen such as liquid nitrogen (T=77 K) orliquid helium (T=4 K), in a preferred embodiment of the invention, therequired cooling is obtained using a closed-cycle cryogenic refrigeratorknown as a cryocooler. Such cryocoolers are available commercially, andcan provide reliable cooling with only a source of electrical powerprovided.

According to a second embodiment, shown in FIG. 2, an externalmultiplexer controller is used to select the signal from each pickupcoil in sequence, with digital processing applied only to the selectedsignal. This has an advantage in reducing the required digital hardware.More generally, one may design a multiplexed system to reduce the Ndigital processing chains to any number between 1 and N, depending on anappropriate compromise of system complexity and performance.

A third embodiment is shown in FIG. 3, in which a digital switch matrixis used as a digital multiplexer that functions similarly to that inFIG. 2. These examples illustrate the flexibility of digital processing,which may be carried out using superconducting circuits in a cryostat,or semiconductor circuits at room temperature, or in an optimizedcombination of the two.

The array of coupling coils may be distributed across an areaapproximately as shown in FIG. 4. This represents a closely spacedtwo-dimensional array, located close to the patient or the tissue to beimaged. Alternatively, a linear array of long pickup coils may be used,as shown in FIG. 5.

It is also possible to provide a three dimensional array of detectors,each having a different spatial sensitivity pattern to magnetic fields,in addition to possible sensitivity to different frequencies.

While some parts of the measurement procedure of applied field gradientsand pulse excitations will be similar to those in conventional MRI,spatial information in one direction (for the linear array) or in twodimensions (for the area array) will be obtained by digital analysis ofthe signals from the respective receiver elements, rather than byrepeated application of magnetic field gradients. This will dramaticallyenhance the speed of the image acquisition.

Each SQUID-based detector in the preferred embodiments may comprise adigital SQUID, similar to that disclosed in U.S. Pat. No. 5,420,586,expressly incorporated by reference. In another preferred embodiment, aSQUID-based detector may comprise a superconducting analog-to-digitalconverter which generates SFQ voltage pulses, similar to that disclosedin U.S. Pat. No. 7,365,663 or U.S. Pat. No. 7,598,897, each of which isincorporated herein by reference. In yet another preferred embodiment,each SQUID-based detector may itself comprise a plurality of SQUIDs, asin a superconducting quantum interference filter (SQIF).

The resonant detection frequency is not limited to low kHz frequenciesand ultra-low magnetic fields, as for conventional analog SQUIDs withexternal flux-locked control loops. For example, consider a magneticfield of 0.2 T, which corresponds to f=8.5 MHz. This field isintermediate between that of conventional high-field MRI andultra-low-field MRI of the prior art. Such a field can be obtained byeither a permanent magnet, conventional electromagnet, orsuperconducting magnet, or combination thereof as may be appropriate. Afield of this magnitude, although too large to be directly exposed tothe SQUID electronics, may be shielded by proper design and location ofthe circuits. The MHz frequency is compatible with digital SQUIDelectronics.

In general, stronger magnetic fields will yield stronger RF signals andbetter imaging resolution. However, the digital SQUID detectors shouldpermit a favorable tradeoff between high resolution and avoidance of thelargest field magnitudes.

In accordance with an aspect of the technology, the excitation andsignal processing need not proceed according to traditional FourierTransform, rectangular coordinate paradigms. Thus, for example, a numberof techniques have been proposed which use different algorithms andtransforms to acquire and present the image data. In particular, the useof broadband superconducting digital SQUID sensors, for example, permitsanalysis of broadband data, and thus relieves a general constraint foundin the art that the received data be relatively narrow band, and fill inonly a very small portion of the k-space matrix per excitation. See,e.g., Dimitris Mitsouras, Frank J. Rybicki, Alan Edelman, Gary P.Zientara, “Fast Magnetic Resonance Imaging via Adaptive BroadbandEncoding of the MR Signal Content”, Magnetic Resonance Imaging, 24;1209-1227, 2006, expressly incorporated herein by reference. Mitsouraset al. propose a wavelet excitation, which may be used in accordancewith the present system. More generally, however, after an excitation, aspread spectrum (or optimized spectrum) probe of all or a portion of thek-space may be conducted, yielding a corresponding modulated/encodedspread spectrum output, which must be faithfully acquired for imagereconstruction. The probe signal is preferably optimized forsignal-to-noise ratio, and desired duration of the scan.

One aspect of the present system is that, in addition to magnetic fieldencoding of space, the use of a plurality of sensors allows directspatial encoding of emissions; that is, n emissions aliased at the samefrequency can be distinguished by using n sensors, and generally ifmultiple excitations are conducted with different probe patterns,further ambiguity may be resolved, and noise reduced. Thus, a singleexcitation pulse can be used to acquire image data from a volume, andrepeated excitation used to increase resolution and reduce noise. Seealso, D. K. Sodickson and W. J. Manning, “Simultaneous Acquisition ofSpatial Harmonics (SMASH): Fast Imaging with Radiofrequency CoilArrays,” Magnetic Resonance in Medicine, vol. 38, pp. 591-603, 1997; K.P. Pruessman, M. Weiger, M. B. Scheidegger, and P. Boesiger, “SENSE:Sensitivity Encoding for Fast MRI,” Magnetic Resonance in Medicine, vol.42, pp. 952-962, 1999; W. E. Kyriakos, L. P. Panych, S. F. Kacher, C.-F.Westin, S. M. Bao, R. V. Mulkern, and F. A. Jolesz, “SensitivityProfiles from an Array of Coils for Encoding and Reconstruction inParallel (SPACERIP),” Magnetic Resonance in Medicine, vol. 44, pp.301-308, 2000; Djaudat Idiyatullin, Curt Corum, Jang-Yeon Park, MichaelGarwood, “Fast and quiet MRI using a swept radiofrequency”, Journal ofMagnetic Resonance 181 (2006) 342-349; Wright, S. M. McDougall, M. P. KeFeng Hollingsworth, N. A. Bosshard, J. C. Chieh-Wei Chang, “HighlyParallel Transmit/Receive Systems for Dynamic MRI”, 31st AnnualInternational Conference of the IEEE EMBS, Minneapolis, Minn., USA, Sep.2-6, 2009, William Scott Hoge, “An Adaptive Signal Processing Approachto Dynamic Magnetic Resonance Imaging”, Ph.D. Dissertation, NortheasternUniversity (2001), each of which is expressly incorporated herein byreference.

Another aspect of the technology is the use of non-rectangularcoordinate systems, and more particularly the permissive use ofnon-linear magnetic gradients. In general, linear gradients are requiredin systems which directly map field strength to position, and thereforethat a magnetic field strength distortion leads directly to an imagedistortion; further, severe distortion leads to aliasing. By using ageneric algorithm that does not presume field linearity, and especiallywith a spatial array of sensors, the requirement of linear magneticfield may be substantially relaxed, to permit, for example,uncompensated magnets or relatively simple configurations of magnets,and permanent magnets. See, for example, David A. Thayer, “ImagingTechniques and Hardware For Inhomogeneous MRI”, Master's Degree Thesis,Department of Electrical and Computer Engineering, Brigham YoungUniversity (2004), expressly incorporated herein by reference in itsentirety.

A further aspect of the technology provides a magneticpolarization-responsive sensor array, which can be used to filter noise,since the noise process has different polarization characteristics fromthe MRI signal process. This polarization filtering extends the sensorarray to extract signal vectors, rather than simple scalar values.

A still further aspect of the technology provides an adaptive imagingtechnique in which an initial estimate of the image is constructed,which for example may be based on a model of the object to be imaged andprior data obtained during the same session. The model may be derivedfrom extrinsic data sources, such as a description of the anatomy, videoobservation, X-ray, ultrasound, or other data. As the model is conformedto the MRI data being obtained, ambiguities may be tested by speciallydesigned excitation to distinguish between multiple possible states.Such a model may also be dynamic, and therefore may be synchronized torespiration, cardiac cycles, neuromuscular activity, and the like. Thus,a movement of a patient need not lead to artifacts, if the degrees offreedom of the patient are understood and accounted for. This adaptivesystem generally requires real-time processing of the data, but usingsuperconducting digital processors, fast traditional semiconductorprocessors, parallel processing computational arrays, such as graphicsprocessor units (GPUs), such processing can be completed between eachexcitation cycle, allowing the image to be resolved in real time.

It is therefore an object to provide a magnetic resonance system,comprising at least one SQUID, configured to receive a radio frequencyelectromagnetic signal, in a circuit configured to produce a pulsatileoutput having a minimum pulse frequency of at least 1 GHz which isanalyzed in a processor with respect to a timebase, to generate adigital signal representing magnetic resonance information.

It is a further object to provide a magnetic resonance method,comprising: receiving a radio frequency electromagnetic signal, in acircuit comprising a SQUID which produces a pulsatile output having aminimum pulse frequency of at least 1 GHz; analyzing the pulsatileoutput with respect to a timebase in an automated processor; andgenerating a digital signal representing magnetic resonance information.

The processor may comprise at least one rapid single flux quantumcircuit.

The magnetic resonance information may comprise magnetic resonance imageinformation derived from an object-under-examination.

The processor may receive a set of outputs from each of a plurality ofSQUIDs, each SQUID responding to a respective radio frequencyelectromagnetic signal from a respective antenna having a respectivelydifferent spatial placement, with respect to theobject-under-examination.

The radio frequency electromagnetic signal may be received from a coilin close proximity to an object-under-examination within a magneticfield, and wherein the magnetic resonance image information is presentedas spatial data.

The processor may be configured to digitally compensate for aninhomogeneous magnetic field.

The processor may be further configured to define a radio frequencyexcitation signal for addressing volumetric regions within a magneticfield.

The processor may defines or generate a broadband radio frequencyexcitation signal for addressing at least one of a spatial range ofvolumetric regions and a plurality of different magnetically resonantisotopes.

The processor may receive the pulsatile output of the SQUID anddecimates in time to produce a signal representation having an updaterate less than the pulse frequency of the pulsatile output and a datarepresentation at each update of greater than 1 bit.

A plurality of antennas may be provided, distributed in space, theantennas each feeding a respective one of a plurality of SQUIDs, eachSQUID forming part of a circuit configured to produce a pulsatile outputhaving a minimum pulse frequency of at least 1 GHz which is analyzed inthe processor with respect to the timebase, to generate a digital signalrepresenting magnetic resonance information from a space proximate tothe antennas.

The SQUID may be configured to represent in its pulsatile output a radiofrequency magnetic field comprising frequencies of at least 100 kHz,whereby the digital signal is substantially oversampled.

The magnetic resonance information may comprise magnetic resonance imageinformation derived from an object-under-examination, further comprisingoutputting an image representation of the magnetic resonance imageinformation. The magnetic resonance information may further comprisemagnetic resonance spectroscopic information, which may be used tocomplement the image representation. The object-under-examination maycomprise a living human body or portion thereof for medical imaging, ora biomedical object or tissue that is subject to high-resolution ormicroscopic examination. Alternatively, the object-under-examination maycomprise a non-biological object that is under examination forspectroscopic or microstructural analysis.

BRIEF DESCRIPTION OF THE FIGURES

FIG. 1 shows a block diagram of a preferred embodiment of the invention,with N parallel digital SQUID-based receivers.

FIG. 2 shows a block diagram of an alternate preferred embodiment of theinvention, where the digital SQUIDs are controlled by a multiplexer thatfeeds the outputs to a single digital processing chain.

FIG. 3 shows a block diagram of a third preferred embodiment of theinvention, where the outputs of the digital SQUIDs are fed to a digitalmultiplexer for sequential readout.

FIG. 4 shows a conceptual picture of a two-dimensional array of pickupcoils, lying above the object to be imaged.

FIG. 5 shows a conceptual picture of a one-dimensional array of pickupcoils, lying above the object to be imaged.

FIG. 6 shows a block diagram of the heterodyne radio receiver for aconventional MRI system.

FIG. 7 shows a block diagram for a direct digital radio receiver using asuperconducting SQUID digitizer, representing an embodiment of thepresent invention.

FIG. 8A shows a conceptual diagram of an MRI system according to oneembodiment of the present invention, using a single pickup coil butseveral distinct resonant frequencies.

FIG. 8B shows the spectral density of transmitted and received RFsignals (and channelized baseband signals) corresponding to the systemof FIG. 8A.

FIG. 9 shows a conceptual block diagram of a two-dimensional MRI antennaarray with a linear array of digital receivers.

DETAILED DESCRIPTION OF THE INVENTION

In a conventional MRI system of the prior art, DC magnetic gradientcoils produce a magnetic field gradient that is scanned across theobject to be imaged. At a given time, only a single volume slice is inresonance with the RF source (and thus excited), and only a single linein K-space is accessed, and the image is developed sequentially. Thisscanned, sequential nature is what makes the imaging so slow.

In a preferred embodiment of the MRI system, as shown in FIG. 4, eachelement of the detector array comprises a magnetic pickup coil orantenna, which is designed to selectively detect local electromagneticfields surrounding the pickup coil, and generally emitted close to thepickup coil. The entire area may be in resonance, so that scanning of agradient field across this area is not necessary, and each pixel isderived from a given antenna element. The parallel processing of thedata from each antenna is what makes the imaging much faster, with atotal imaging time that may ultimately approach the pulse relaxationtimes less than 1 s. This may permit imaging of images moving orchanging in time, i.e., video imaging.

It may still be necessary to apply a gradient field in the thirddimension, corresponding to selecting a slice parallel to the array,into the depth of the object. In an alternate preferred embodiment ofthe invention shown in FIG. 5, the detector array may comprise a set oflong narrow parallel pickup coils arrayed along a single direction. Inthat case, one would need a gradient field to excite a line parallel tothe array direction (and perpendicular to the coil length) to providespatial information perpendicular to the array direction. Such aresonant line could be scanned across the object, as well as through itsdepth. Clearly, imaging using a resonant area would proceed faster thana resonant line, which would be faster than a resonant voxel. Ingeneral, the greater parallelism requires a greater number of receiverelements. The balance between speed and system complexity would bedetermined by the needs of a given application.

In greater detail, each coil may comprise an inductive coil withinductance L (which may have multiple turns), designed to detect the RFmagnetic field of the signal emitted by the object. Each coil may alsobe a resonant coil at the detection frequency, whereby a capacitor C iscombined with the inductor corresponding to an LC resonator such thatthe resonant frequency f=½π√(LC) is the desired detection frequency. Theinductor may comprise a superconducting inductor, which will tend toincrease the quality factor Q of the resonator. A higher Q is generallypreferable, provided the bandwidth is large enough to measure the entireRF signal; a higher Q receiver would receive less broadband noise.

In a preferred embodiment, a pickup coil may comprise a first-derivativegradiometer or a second-derivative gradiometer, as disclosed, forexample, in U.S. Pat. No. 7,002,341, “Superconducting quantuminterference apparatus and method for high-resolution imaging ofsamples”, expressly incorporated herein by reference. Such an RFgradiometer coil (which is to be distinguished from the DC gradientfield coils) comprises a compound inductor designed to cancel uniformmagnetic fields (and uniform field gradients for the second-derivativecase). In this way, a gradiometer coil is far more sensitive to signalsemitted from sources very close to the coil, rather than sources furtheraway. This permits one to directly obtain spatial resolution from eachreceiver coil. The spatial resolution from the coils will be used incombination with the resonant volume, area or line to provide imaging inthree dimensions.

Note that a gradiometer signal may alternatively be obtained bysubtracting signals from adjacent pickup coils further in the dataprocessing. However, gradiometer coupling at the front end shouldenhance the effective dynamic range of the detectors.

The data processing chain for each receiver is shown in FIG. 1. The RFsignal from the pickup coil is coupled to a digital SQUID, whichgenerates single-flux-quantum (SFQ) digital pulses at a high data rate(typically of order 10 GHz or greater). The RF signal is a narrow-bandsignal at f=γB (where B is the measurement magnetic field and γ=43MHz/T) which may be from the kHz to the MHz range. As described above, adigital SQUID can measure RF fields well into the MHz range, unlike aconventional analog SQUID with an external control loop which is limitedto a few kHz. Since the GHz data rate from the SQUIDs is much higherthan the MHz magnetic signal to be analyzed, this represents anoversampled digital signal. The higher frequency MHz range maycorrespond to stronger signals which may provide higher-resolutionimages.

The required signal for imaging is actually a relaxation time of the RFpulse after excitation, typically of order 0.1-1 s. (There are severaldistinct relaxation times, referred to in the literature as T₁, T₂ andT₂*.) One can regard the slow relaxation as a baseband signal thatmodulates the RF carrier. So it is useful to downconvert the RF signaland extract this baseband signal digitally, using a digital localoscillator. The resulting signal can be digitally averaged using adigital decimation filter (effectively a binary counter) to increase thenumber of bits and decrease the bandwidth. Some or all of this digitalprocessing may be carried out using superconductingrapid-single-flux-quantum (RSFQ) electronics, which is matched to theoutput format of the digital SQUID. The digital baseband signal can thenbe amplified and sent out of the cryostat to interface with conventionalsemiconductor digital electronics at room temperature for furtherdigital processing and image generation.

While the digital baseband signals from each of the digital receiverscould in principle be sent out in parallel, it may be advisable todecrease the number of data lines coming out of the cryostat. A largenumber of such data lines may conduct heat into the cryostat, which isundesirable. One type of data line reduction is serialization, whereby nbits are sent out sequentially at a higher data rate. In addition, the Nsignals from the N receivers could be digitally multiplexed. Thedemultiplexing and deserialization can be carried out using conventionalsemiconductor digital electronics at room temperature.

Two other techniques for digital multiplexing are illustrated in FIGS. 2and 3. FIG. 2 shows an external multiplexer controller, which mayprovide external power to activate each of the SQUIDs in sequence. Thisis similar to the time-domain SQUID multiplexing that has beendemonstrated in the prior art for arrays of analog SQUID amplifiers forcryogenic sensor arrays. (See, for example, “Superconducting multiplexerfor arrays of transition-edge sensors, J. Chervenak et al., AppliedPhysics Letters, vol. 74, p. 4043, 1999.) Alternatively, one may use afully digital multiplexer similar to that in FIG. 1, but one thatoperates at higher frequencies. (See, for example, “Superconductingdigital multiplexers for sensor arrays”, A. Kadin et al., ProceedingsThermal Detectors Workshop, NASA, 2003.) Either of these schemes has anadvantage in reducing hardware duplication, resulting in a more compactdigital processor at cryogenic temperatures, while maintaining theparallel processing with accelerated imaging rate.

Superconducting devices must be cooled to cryogenic temperatures forproper operation. At present, the most widespread digitalsuperconducting electronics technology is comprised of niobium (Nb)Josephson junctions, which can operate below the critical temperaturebelow 9 K, and generally are operated below 5 K. These may be installedin a cryostat, which may be cooled either by liquid helium, or using amulti-stage cryocooler. Alternatively, high-temperature superconductorssuch as YBa₂Cu₃O₇ (YBCO) may be used, with a critical temperature of 90K. Such a system may operate in liquid nitrogen (at 77 K), or with asingle-stage cryocooler at temperatures of 40 K or above. While thereliability and performance of YBCO SQUIDs and digital electronics arecurrently inferior to that of Nb, the same circuit architectures may beapplied if and when these or other higher-temperature materials becomepractical.

It is necessary to design the pickup coil assembly to lie close to theobject to be imaged, without subjecting the object (which may be a humanpatient) to cold temperatures. This requires that the superconductingdevices be properly packaged inside a cryostat with vacuum jacketing.Further, the RF signals must pass through the cryostat walls withoutloss, so that metallic jackets and shields cannot be used. Such anon-metallic cryostat has been demonstrated in the prior art, usingcomponents such as reinforced fiberglass.

It may also be necessary to shield the Digital SQUIDs andsuperconducting electronics from a large magnetic field that may be usedas a polarizing field or a measuring field. While the pickup coils mustbe near the object to be imaged, and thereby close to the large magneticfield, the superconducting devices can be located inside a magneticshield, which may be some distance away from the peak magnetic field.Appropriate magnetic shield materials may include superconducting layersas well as soft ferromagnetic materials such as mu-metal. The pickupcoils may be spread over a relatively large area, but thesuperconducting devices may be concentrated on a small number of chipson a multi-chip module located in a central, shielded assembly.

Note that the term “Digital SQUID” in FIGS. 1-3 refers to anysuperconducting device comprised of Josephson junctions that convertsmagnetic flux to digital pulses, e.g., SFQ digital voltage pulses.Several such circuits are described in the review article on“Superconducting analog-to-digital converters”, O. Mukhanov et al.,Proceedings of the IEEE, vol. 92, p. 1564, 2004.

In another preferred embodiment, the array of coupling coils as in FIGS.4 and 5, and analogously, a volumetric array, may be scanned across theobject to be imaged (or the array held fixed and the object moved). Thiswould permit imaging of a larger object to high spatial resolution,without requiring a proportionally larger number of array elements.While this scanning would slow down the imaging process, theacceleration permitted by the array parallelism may make this practical.

In yet another preferred embodiment, an additional mode of parallelismmay be associated with the RF excitation signal. For example, one mayapply a deliberate magnetic field gradient such that one plane isselected to have resonant frequency f₁ and another adjacent plane tohave resonant frequency f₂. If the RF excitation signal (from one ormore transmit antennas) simultaneously comprises appropriate pulses withfrequencies f₁ and f₂, then the RF decay signal will comprise componentsat both frequencies. If both of these frequencies are within thebandwidth of the digital SQUID detector, then both signals will bedetected, but can be separated by subsequent digital filtering or othertypes of analysis. This provides an example of frequency-domainmultiplexing, with potential processing speedup proportional to thenumber of frequencies N selected, which are clearly not limited to two.For a two-dimensional array such as that in FIG. 4, this approach wouldpermit simultaneous selection of N parallel slices in resonance.

The main thrust of this technology is to provide parallel processing toenable fast imaging, at rates that may be faster than pulse rates orbreathing rates, or functional MRI with a single stimulus. However, themassively parallel processing may also enable other approaches to MRIthat are conventionally too slow. For example, while MRI generally usesthe proton signal (from hydrogen in water and organic compounds), otheratomic nuclei such as isotopes of Na and P also exhibit magneticresonance, with a much weaker signal due to the lower concentration ofthese atoms. Extensive signal averaging or other extended temporalsignal processing, would be useful to obtain a high-resolution image,but the speed-up and low-noise detectors provided herein may make thisfeasible.

As described above, MRI is conventionally based on a narrow-band radiocommunications system, with a narrow-band transmit signal and anarrow-band receive signal, where the frequency is proportional to avalue of magnetic field. The bandwidth of the receive signal istypically less than 100 kHz, for a radio signal that may be typically inthe range from 40 MHz to 130 MHz. For this reason, a conventionalheterodyne receiver is typically used for MRI, as shown in FIG. 6, withan antenna followed by a low-noise amplifier, an analog mixer todownconvert the signal to a lower frequency, and a receiver for thedownconverted baseband signal. In modern MRI receivers, a digitalbaseband receiver is used, with an analog-to-digital converter thatoperates on the baseband signal, producing a digital signal that can beused to process the image. The sampling rate of this baseband ADC neednot be more than about 1 MHz.

In contrast, in the simplest corresponding system of an embodiment ofthe technology, shown in FIG. 7, a wideband direct digital RF receiveris used, instead of the heterodyne receiver of FIG. 6. In particular, awideband superconductor ADC is used, which has a sampling frequency thatis in excess of 1 GHz, which may be 20 GHz or higher. For an RF signalat 100 MHz, this is extreme oversampling, which might normally be viewedas unnecessary for this application. Indeed, for a single narrow-bandsignal, such a receiver is unnecessary and not well matched to theapplication. However, one can present a direct analogy with a modernmulti-user communication system, which increasingly makes use ofbroadband receivers to simultaneously receive a wide band comprising aplurality of narrow-band signals. If an MRI system is extended tomultiple signals that are multiplexed in the frequency, time, and codedomains, then a broadband receiver will make more efficient use of theavailable spectrum with a minimum of hardware. An additionalconsideration for MRI is that scans are generally quite slow, andparallelizing the component signals in time and/or frequency will enablefaster scans.

It is notable that the wideband superconductor ADCs as typicallyemployed herein, are essentially digital SQUIDs, with the sensitivityand low noise that implies. The required gain for the low-noiseamplifier may be substantially reduced, or in some cases the LNA may beeliminated entirely, with the pickup coil within the magnetic field, andthe SQUID shielded from the high magnetic field but in close proximity,or the SQUID separated from the pickup coil by, e.g., a coaxial cable,with a low noise amplifier (LNA) used to transmit the signal.Conventional MRI receivers typically use this split receiver approach.The coil and LNA may be cooled, for example by a compact 70K cryocooler,to reduce their noise, and a separate 4K cryocooler provided for thelow-T_(c) superconducting circuits in the adjacent instrument room.

In one embodiment, a plurality of radio frequencies are simultaneouslyexcited, corresponding to different slices in the body being examined(FIG. 8A). A single wideband receiver can be used to receive all ofthese frequencies simultaneously, and they can be separated using adigitally channelizer (FIG. 8B). This parallel processing can lead tosome speedup in generation of 3D images. This simultaneous multi-sliceapproach was described in the prior literature (see, e.g., J. H. Weaver,“Simultaneous multislice acquisition of MR images”, Magnetic Resonancein Medicine, vol. 8, pp. 275-284, 1988), expressly incorporated hereinby reference, and demonstrated for a small number of frequencies, butnot implemented in practice because of the lack of an appropriatebroadband receiver. See also, US 2009/0278538, expressly incorporatedherein by reference.

In an alternative embodiment, a plurality of pickup coils or antennasmay be used. These may be arrayed as surface coils along the surface ofthe object to be imaged, as shown in FIGS. 4 and 5. These might comprisea 1D array of coils that can be scanned, a 2D array of coils (FIG. 4),or a 1D array of long coils (FIG. 5). In one configuration, each suchpickup coil may be connected a separate receiver. While such parallelcoil arrays are being implemented using conventional technology, themultiplication of analog RF receivers and processing arrays, eachindependently calibrated, has problems with scaling to large arrays.

In one embodiment, the hardware for multiple coils is simplified byusing direct digital receivers with digital signals that may bemultiplexed. For example, FIG. 1 shows an array of broadband directdigital receivers, each based on an oversampled superconducting ADC(also described as a digital SQUID). The narrowband signals aredigitally extracted, and digitally combined in time, frequency, or codedomains using a high-speed digital multiplexer. Variants of this areshown in FIGS. 2 and 3, where the digital multiplexer is applied to thebroadband digital signal earlier in the processing chain. FIG. 2 shows asystem which employs time-domain multiplexing of signals with anexternal controller. FIG. 3 shows a more general multiplexer forcombining digital signals.

In another preferred embodiment, one may have a plurality of directdigital receivers, each of which combines the inputs from a plurality ofcoupling coils. For example, in FIG. 9, the signals from a row ofcoupling coils (which are assumed to represent signals that areappropriately orthogonal in frequency, time, or code) are combined onthe same transmission line that feeds a SQUID ADC (oversampledsuperconducting ADC). But there are also multiple rows. Together, theseenable spatial information in 2 dimensions. These can be combined withconventional resonant excitation from a transmit signal to obtainspatial resolution in the z-direction.

In principle, the superconductor MRI system could apply to systems witheither large magnetic fields or small magnetic fields, with frequenciesfrom 1 MHz to 500 MHz or more. Large fields provide larger signals,higher signal to noise ratio, and that is the direction that thetechnology is moving. But large fields are expensive and heavy, andcreate problems with RF heating, acoustic noise, and issues of safetyand imaging artifacts. If one could obtain the same imaging speed andresolution with a 0.5 T system as with a 1.5 T system, the lower fieldwould be preferred. The superconductor digital-SQUID receivers should bemore sensitive than conventional receivers, particularly for relativelylow frequencies, permitting operation in different regimes thantraditional sensors.

The embodiments presented here are not exclusive, but are used toillustrate the wide range of flexible digital processing solutions thatare enabled by the use of broadband digital receivers.

There has thus been shown and described detector methods and systems formagnetic resonance imaging which fulfill all the objects and advantagessought therefor. Many changes, modifications, variations, combinations,subcombinations and other uses and applications of the subject inventionwill, however, become apparent to those skilled in the art afterconsidering this specification and the accompanying drawings whichdisclose the preferred embodiments thereof. All such changes,modifications, variations and other uses and applications which do notdepart from the spirit and scope of the invention are deemed to becovered by the invention, which is to be limited only by the claimswhich follow.

What is claimed is:
 1. A magnetic resonance imaging system, comprising aplurality of SQUIDs, each being configured to receive a radio frequencyelectromagnetic signal, and provided within a circuit configured toproduce a pulsatile output having a minimum pulse frequency of at least1 GHz which is analyzed in a digital processor with respect to atimebase, to generate a digital signal representing magnetic resonanceimage information derived from an object under examination, wherein theprocessor receives a set of outputs from each of a plurality of SQUIDs,each respective SQUID responds to a respective antenna having arespectively different spatial placement, with respect to theobject-under-examination.
 2. The magnetic resonance imaging systemaccording to claim 1, wherein the digital processor comprises at leastone rapid single flux quantum circuit.
 3. The magnetic resonance imagingsystem according to claim 1, wherein the radio frequency electromagneticsignal is received from a coil in close proximity to anobject-under-examination within a magnetic field, and wherein themagnetic resonance image information is presented as spatial data. 4.The magnetic resonance imaging system according to claim 1, wherein thedigital processor is configured to digitally compensate for aninhomogeneous magnetic field.
 5. The magnetic resonance imaging systemaccording to claim 1, wherein the digital processor is furtherconfigured to define a radio frequency excitation signal for addressingvolumetric regions within a magnetic field.
 6. The magnetic resonanceimaging system according to claim 1, wherein the digital processordefines a broadband radio frequency excitation signal for addressing atleast one of a spatial range of volumetric regions and a plurality ofdifferent magnetically resonant isotopes.
 7. The magnetic resonanceimaging system according to claim 1, wherein the digital processorreceives the pulsatile output of the plurality of SQUIDs and decimatesin time to produce a signal representation of each respective pulsatileoutput having an update rate less than the pulse frequency of therespective pulsatile output and a data representation at each update ofgreater than 1 bit.
 8. A magnetic resonance digital system, comprising:at plurality of SQUIDs, each being configured to receive a radiofrequency electromagnetic signal, and being provided in a respectivecircuit configured to produce a pulsatile output having a minimum pulsefrequency of at least 1 GHz which is analyzed in a digital processorwith respect to a timebase, to generate a respective digital signalrepresenting magnetic resonance information; with respect to anobject-under-examination; and a plurality of antennas, distributed inspace, the antennas each feeding a respective one of the plurality ofSQUIDs, each respective digital signal representing magnetic resonanceinformation from a space proximate to a respective antenna.
 9. Themagnetic resonance imaging system according to claim 1, wherein each ofthe plurality of SQUIDs is configured to represent in its pulsatileoutput a radio frequency magnetic field comprising frequencies of atleast 100 kHz.
 10. The magnetic resonance imaging system according toclaim 1, wherein signals from a plurality of antennas are multiplexedand presented to a common SQUID.
 11. The magnetic resonance imagingsystem according to claim 1, wherein signals from the plurality ofSQUIDs are multiplexed into a composite digital signal, presented foranalysis by the digital processor.
 12. A magnetic resonance digitalmethod, comprising: providing a plurality of SQUIDs, each SQUIDresponding to a radio frequency electromagnetic signal from a respectiveantenna having a respectively different spatial placement, with respectto an object-under-examination, and being disposed in a circuit whichproduces a pulsatile output having a minimum pulse frequency of at least1 GHz; analyzing the pulsatile output of each respective SQUID withrespect to a timebase in an automated digital processor; and generatinga digital signal representing magnetic resonance image informationderived from the object under examination, wherein the automated digitalprocessor receives a set of outputs from each of a plurality of SQUIDs.13. The magnetic resonance digital method according to claim 12, whereinthe automated digital processor comprises at least one rapid single fluxquantum circuit.
 14. The magnetic resonance digital method according toclaim 12, further comprising outputting an image representation of themagnetic resonance image information.
 15. The magnetic resonance digitalmethod according to claim 12, wherein the analyzing further comprisingdigitally compensating for an inhomogeneous magnetic field.
 16. Themagnetic resonance digital method according to claim 12, wherein thegenerating further comprising generating a radio frequency excitationsignal for addressing volumetric regions within a magnetic field. 17.The magnetic resonance imaging method according to claim 12, wherein thegenerating further comprising generating a broadband radio frequencyexcitation signal.
 18. The magnetic resonance imaging method accordingto claim 12, wherein the generating further comprising decimating intime to produce a signal representation having an update rate less thanthe pulse frequency of the pulsatile output and a data representation ateach update of greater than 1 bit.
 19. The magnetic resonance imagingmethod according to claim 12, wherein the SQUID is configured torepresent in its pulsatile output a radio frequency magnetic fieldcomprising frequencies of at least 100 kHz.
 20. A magnetic resonanceimaging method, comprising: receiving a radio frequency electromagneticsignal, in a circuit comprising a plurality of SQUIDs, each of whichproduces a pulsatile output having a minimum pulse frequency of at least1 GHz; analyzing the pulsatile output of each respective SQUID withrespect to a timebase in an automated digital processor; generating adigital signal representing magnetic resonance image information; andproviding a plurality of antennas, distributed in space to receivesignals from an object-under-examination, the antennas each feeding arespective one of the plurality of SQUIDs, each SQUID forming part of acircuit configured to produce the pulsatile output having a minimumpulse frequency of at least 1 GHz which is analyzed in the automateddigital processor with respect to the timebase, to generate a digitalsignal representing magnetic resonance information from a spaceproximate to the antennas.